There are schematically three types of resorbable biomaterials. The first one is a biomaterial which occupies a void or a virtual space and is applied for the occlusion of vessels or other cavities (natural or surgical ones), defects such as wrinkles. The second type is a biomaterial which has a pure function as a drug delivery system examples: local delivery in organs such as brain (antimitotic agent), eye (antiangiogenic agent), or cavities (antibiotic or anti-inflammatory agent in surgical voids). The third type is a biomaterial which combines the function of a space filler with a delivery function (embolization microsphere delivering an antimitotic agent, dermal filler containing an anesthesic or anti-inflammatory drug).
The two functions which are required for resorbable drug delivering biomaterials are not fully assumed by the existing resorbable biomaterials. Although they have individually some interesting properties, they don't possess enough of these properties to be proposed as a multifunction implantable biomaterial. An ideal material should swell in a controlled way in situ after implantation, deliver a drug in controlled manner in terms of time and rate and finally resorb after its delivery. The following examples, drawn from the field of embolization and tissue bulking, are given to illustrate the insufficiencies of the existing resorbable materials.
In the field of embolization, several products have individually one interesting property. Gelatin sponges are biodegradable after implantation in tissues or injection in cavities, ducts or vessels. They can easily be impregnated with physiological saline and/or contrast media. However, after their hydration they loose their shape and resistance. In addition, there is a great variability in resorption speed, which is influenced by many factors such as nature, homogeneity, size, enzymatic potential, and local inflammatory response. Moreover, since the mass of resorbable gelatin may vary in large proportions, the resorption time of the plug will consequently also take a variable time.
Dextran starch microspheres (Spherex® from Pharmacia; Embocept® from Pharmacept) are non-toxic, readily degradable and notably used to provide temporary vascular occlusion, mainly for the treatment of tumor when co-administered with chemotherapeutic drugs. However, they suffer from several limitations. First of all, these microspheres are available only in small sizes, with diameters below 100 μm. Such a small diameter does not allow targeted embolization, particularly for proximal occlusion. Besides, resorption is fast, with a usual half life below 1 hour, and cannot be accurately predicted since depends on the enzymatic capability to resorb a given microspheres volume.
Water-absorbent dry microspheres based on acrylic and PVA copolymers have also been proposed as swellable implants for embolization (Osuga et al. (2002) J Vasc Intery Radiol. 13:929-34). In a commercial presentation (Quadrasphere®, Biosphere Medical), these microspheres are under a dry form. For their use they are mixed with physiological saline, and/or iodinated contrast media. Compared to their initial size, their final size after water uptake varies according to the ionic charge of the medium (×2 or ×4 in saline and contrast medium respectively). However the final size varies too much to allow for their controlled final volume after implantation, which is a serious limitation for their use. Besides, these microspheres are not resorbable.
In the field of soft tissue repair and augmentation, a number of products have been used. However, they present all some disadvantages:
Silicone gel (or silicone oil) is easy to use. However, the migration of droplets of silicone into the tissues situated below the point of injection, by simple gravity, has been observed after injection. It appears also that liquid silicone tends to migrate to distant body part causing a variety of physiological and clinical problems. Indeed, silicone is frequently the cause of chronic inflammation, of formation of granulomas, and even of tardive allergic reactions. Silicone is not biodegradable, and it is often found in the liver. Therefore, the FDA has prohibited the use of liquid silicone in humans.
Collagen suspensions have been very widely used in the last ten years. The results have however been quite disappointing since collagen is resorbed within 1 to 3 months. It should also be noted that collagen is of bovine origin and allergic reactions to the bovine proteins are noted in about 2% of patients. In an attempt to solve these problems, crosslinked collagen was introduced to extend effective treatment times to approximately six months. However allergic reactions still occur.
Hyaluronate gels provided a good alternative by virtue of their biocompatibility and their lack of toxicity. They are moreover widely used in eye surgery. However, their rapid bioresorbability (maximum 2 months) makes them ineffective for use in plastic surgery. Furthermore, hyaluronic gels can be source of acute or delayed hypersensitivity and can generate severe local inflammatory response.
Particles which are either biodegradable (PLGA) or not (acrylamide, PMMA, EMA) can also be used.
Non biodegradable particles such as poly(methyl methacrylate) (PMMA) microspheres are permanent. Because of that, the body can mount a foreign body response to these polymers and forms a tight fibrous capsule around the material. Furthermore there is a risk of migration of this material away from the injection site.
The disadvantages of biodegradable particles such as PLGA are their tendency to aggregate prior to and/or during clinical application which will render difficult their injection and/or form hard, insoluble nodules at the injection site, causing oedema and swelling and, most of the times, requiring corrective medical intervention. Furthermore they undergo a prolonged inflammatory response as long as the degradation takes place and subunits are released.
It is also known to use a combination of gel material (hyaluronate gel and/or collagen gel) containing microparticles (degradable or not). In particular known commercial products are New-Fill Sculptra® from Sanofi Aventis (poly-L-lactic acid microparticles suspended in sodium carboxymethylcellulose, mannitol and water) and Artefill®, Artecoll® from Artes medical (poly(methyl methacrylate) microspheres suspended in collagen gel). However, the combination of gel and microparticles do not solve the above mentioned problems. The carrier gel disappears from the site within 1 to 3 months and at the same time the host response to the remaining microparticles gradually makes up for the loss of filling effect in a more permanent manner. The host foreign-body response runs its course and ends up until a permanent de novo fibrous scar tissue is embedding the intended filler agent of these gels.
Furthermore, literature has since many years established, for solid implants, that tissue ingrowths in the implanted biomaterial depend in a large part on the porosity of the material. Scaffolds and/or matrix with controlled porosity are required to allow cell ingrowth, nutrient diffusion and sufficient formation of vascular networks. Mean pore size is an essential aspect of scaffolds for tissue-engineering. If pores are too small cells cannot migrate in towards the center of the construct limiting the diffusion of nutrients and removal of waste products. Conversely, if pores are too large there is a decrease in specific surface area available limiting cell attachment. The permeability of scaffolds and other three-dimensional constructs used for tissue engineering applications is important as it controls the diffusion of nutrients in and waste out of the scaffold as well as influencing the pressure fields within the construct (O'Brien Technol Health Care. 2007; 15(1):3-17).
To facilitate the injection of the particles in needles having small diameter, several dermal fillers contain a natural polymer gel, which is resorbed quickly after the implantation. Their resorbtion time is usually homogeneous and occurs generally quickly. Since these gels represent the major part of the injected volume (80% in some cases), it leads the gels to lose a large part of their filler effect from disappearance. This component is said to favor the tissue ingrowth. However since they contain few matters and have a high water content (about 90%) these gels offer to the body a structure, which is often too loose to constitute a matrix for tissue ingrowth. Therefore, most combination gels (gel and particles) are not actually efficient to facilitate the tissue ingrowth between the particles or between the polymer threads.
Microspheres have been proposed to prepare solid scaffolds for tissue engineering by Brown (Brown J Biomed Mater Res B Appl Biomater. 2008 August; 86B(2):396-406). He has applied a technique of solvent/non-solvent sintering which creates porous polymeric microsphere scaffolds suitable for tissue engineering purposes with control over the resulting porosity, average pore diameter, and mechanical properties. Five different biodegradable biocompatible polyphosphazenes exhibiting glass transition temperatures from −8 to 41° C. and poly (lactide-co-glycolide), (PLGA) a degradable polymer used in a number of biomedical settings, were examined to study the versatility of the process and benchmark the process to heat sintering. Parameters such as: solvent/non-solvent sintering, solution composition and submersion time affect the sintering process. PLGA microsphere scaffolds fabricated with solvent/non-solvent sintering exhibited an interconnected porosity and pore size of 31.9% and 179.1 micrometers, respectively which was analogous to that of conventional heat sintered PLGA microsphere scaffolds. Biodegradable polyphosphazene microsphere scaffolds exhibited a maximum interconnected porosity of 37.6% and a maximum compressive modulus of 94.3 MPa. Solvent/non-solvent sintering is an effective strategy for sintering polymeric microspheres, with a broad spectrum of glass transition temperatures, under ambient conditions making it an excellent fabrication route for developing tissue engineering scaffolds and drug delivery vehicles.
The patent WO 2009049230 describes a solid scaffold including a plurality of biocompatible microspheres sets linked together by a partial melding of the microspheres in a solvent or solvent system gaseous sub-critical CO2 to form a three-dimensional matrix. The matrix's pores, defined by and disposed between the microspheres, range from about 200 micrometers to about 1650 micrometers. The different sets of microspheres can have different characteristics, such as polymer nature, particle size, particle size distribution, type of bioactive agent, type of bioactive agent combination, bioactive agent concentration, amount of bioactive agent, rate of bioactive agent release, mechanical strength, flexibility, rigidity, color, radiotranslucency, radiopaqueness. A type of microspheres can be made from a biodegradable polymer such as poly-lactide-co-glycolide or poly(lactic-co-glycolic acid).
However the diameter of the porous spaces located between the microspheres in a sediment of microsphere having a compacity of 60% is rather small, about 13% of the microspheres diameter. It means that in a cluster of microspheres having a mean diameter of 100 μm, the size of the inter-microsphere pore is about 13 μm. If one consider that the clusters resulting from an injection of a microspheres suspension have a lower compacity than 60%, one could consider that in such clusters the pore diameter ranges from 13 μm to a few dozens of micrometers. Such an inter-microsphere pore size is not favourable to tissue ingrowth in the clusters. The minimum pore diameter for one cell like a macrophage is about 15 μm. Moreover some cells like preadipocytes enlarge during differentiation due to incorporation of lipids and then need large pores sizing at least 40 μm (Von Heimburg, Biomaterials. 2001 March; 22(5):429-38).
The growth of an organised tissue in a scaffold needs both large pore size and large void fraction. Void Fraction (VF) of microspheres sediments, which ranges from 60% to 70% is insufficient to allow an efficient filling up by a functional tissue. Zeltinger had demonstrated that Void Fraction, independently from pore size, was a major determinant of scaffold colonisation. Scaffold with 75% VF were unsuitable for tissue formation while those with 90% VF were suitable for tissue formation when pore size was over 38 μm (Zeltinger Tissue Eng. 2001 October; 7(5):557-72). In summary the natural microspheres sediments/beds obtained by microspheres injections in tissues are very far from offering to tissue ingrowth a convenient pore size and a sufficient void fraction.
Therefore there is a need to find a new product for soft tissue repair and augmentation useful as scaffolds and/or matrix for tissue ingrowth.
It has been proposed in patent No PCT/EP 2010/063227 to synthesize swellable and bio-resorbable cross-linked polymers liable to be implanted. However after degradation of these polymers the residues presents in the body may have a too high molecular weight. Therefore, they tend to accumulate in the kidney of the patient which could be detrimental for its health.
Furthermore, when the polymer is loaded with drugs it is important to be able to control over time the rate of release of said drug. The controlled release of the drug is achieved by diffusion, swelling of the network or degradation of the polymeric matrix. Drug loading can be done by chemical conjugation or physical entrapment, and so the respective drug release processes are dependent on the type of encapsulation.
In the chemically conjugated drug, the loading is predictable since it is directly connected to the chemistry of the particles. The release should occur by hydrolysis of the linker between the polymer network and the drug. To control the release, one could play on the chemical composition of the linker. However, it has been observed that for hydrophobic drugs, the degradation of the polymer network could be slow down. Indeed, apolar structure of the drug may decrease the swelling of polymer and so less water molecules will be in contact with the hydrolysable groups of the polymer. Moreover, for the same reason, hydrophobic structure inside the polymer will considerably reduce the release.
For the physically entrapped drug, the loading could be done either by passive adsorption (swelling of the polymer into a drug solution) or by ionic interaction. Efficiency of encapsulation depends mainly on the compatibility between both structures and/or favourable interactions. Generally, polymer prepared using the PLGA system exhibits release kinetics based on both erosion and diffusion. In this type of system, an initial burst or rapid release of drug is observed. This burst effect can result in unwanted side effects in patients to whom the polymers have been administered. To obtain a controlled and/or sustained release of the drugs with crosslinked microspheres of polymer, two main factors may affect drug release: nature and extent of cross-linking. For low molecular weight drugs (typically below 1000 Da), it is not easy to play with these two factors since decreasing the mesh size of the polymer network, will not or poorly modify the diffusion of molecules through the meshes before degradation of the polymer. Like the covalently conjugate system, entrapment of hydrophobic drugs reduces the swelling of the polymer and may change the degradation rate of the resorbable system.
It is therefore a goal of the present invention to solve the above mentioned problems.